1. Field of the Invention
The present invention is broadly concerned with improved implantable biosensors and methods of production thereof. More particularly, the invention pertains to biosensors having enhanced sensitivity and stability characteristics coupled with relative ease of manufacture. The biosensors are preferably fabricated using enzymes, and especially oxidase enzymes such as glucose oxidase.
2. Description of the Prior Art
The enzyme electrode, and especially the glucose electroenzymatic biosensor, has served for more than 25 years as a valuable clinical tool for detecting and monitoring diabetes. A majority of glucose sensors, especially those used in in-vivo applications, are based on the rate of glucose oxidase-catalyzed oxidation of glucose by dioxygen, where the rate of the reaction is measured by monitoring the formation of hydrogen peroxide or the consumption of oxygen. The fabrication of such a sensor involves the controlled deposition of a permselective polymer layer used to eliminate interferences such as ascorbate, urate and acetaminophen, an enzyme layer, and an outer layer that renders the sensor response mass transfer rather than kinetically controlled and which also provides a biocompatible interface with the surrounding environment. The use of thick film techniques, including screen printing, has been demonstrated to be successful in the preparation of sensors with reasonably reproducible characteristics, and this approach has been applied, for example, in the electrochemically-based sensors used for self-monitoring of blood glucose as marketed by Abbott Laboratories (Medisense) and others. If, however, it is desired to employ a cylindrical geometry or to prepare a sensor array, then the reproducible deposition of the various functional layers becomes significantly more complicated. Thus it would be of considerable advantage to control the preparation of the sensor electrochemically, especially when the sensing elements in an array are themselves electrochemically addressable. This would also allow for the deposition of different enzymes in various parts of the array.
Electropolymerization makes it possible to generate a coating on small electrodes of complex geometry and to do so precisely in one or two rapid and simple steps. In general, electrochemically-mediated fabrication of biosensors has been accomplished in two ways. First, a polymer layer is formed directly on the electrode, and polymers formed from such monomers as pyrrole, aniline, tyramine, o-aminophenol and o-phenylenediamine have been used to create a permselective layer before or after the application of enzyme solution and cross-linking with glutaraldehyde. A second approach involves the entrapment of enzyme in a growing polymer network by co-polymerization of enzyme and monomer. In some cases a monomer unit is attached to the enzyme to facilitate this process. Yacynych employed a copolymer of 1,3-diaminobenzene and resorcinol as the preferred film for blocking interferences from the surface of carbon or partially platinized carbon electrodes (Geise, R. J.; et al., Biosens. Bioelectron., 1991, 6, 151-160). Vadgama and coworkers found that electropolymerized 4-aminophenol and then phenol constituted an exceptionally selective film against acetaminophen and ascorbate in glucose biosensors (Eddy, S.; et al., Biosens. Bioelectron., 1995, 10, 831-839). Curulli et al (Carelli, I.; et al., Electrochim. Acta, 1996, 41, 1793-1800 reported the results that poly(1,3-diaminobenzene/catechol) was the most efficient polymer to prevent the interference of acetaminophen. It has also been pointed out that electropolymerized films have significantly different characteristics when formed on different electrode materials and under different electropolymerization conditions. Experience has shown that these approaches typically give sensors of moderate activity but often high selectivity, but that both of these essential characteristics deteriorate quickly over a period of several days. The fact that the diffusion of enzyme and monomer cannot proceed at the same rate makes it difficult to enrich the composite layer with enzyme without at the same time degrading the permselective properties of the polymer.
U.S. Pat. Nos. 5,540,828, 5,286,364, 5,165,407, 5,310,469, 5,411,647, 5,166,063 and 4,721,677 describe various types of electrochemical biosensors.
The present invention overcomes many of the problems outlined above and provides biosensors such as implantable glucose sensors which can be economically prepared and which have excellent selectivity and stability characteristics far in excess of typical prior art sensors. Selectivity, or the ability to exclude electroactive interferants, is defined as, in the case of a glucose sensor, the percent change in the electrical signal at 5 mM glucose according to the relationship % Interf=((ITotxe2x88x92IGlu)/IGlu)xc3x97100, where IGlu is the current response for glucose and ITot is the current response of the glucose and interferant. Current response is defined as the current produced by the sensor in response to the analyte. The current response produced by the sensor in response to the analyte divided by the analyte concentration is herein referred to as the sensor sensitivity. For a biosensor prepared in accordance with this invention, the ratio of the current response at 5 mM glucose to that for 0.1 mM acetaminophen should be no less than about 100 in vitro.
Broadly speaking, a method of preparing a biosensor in accordance with the invention includes providing an electrically conductive biosensor electrode, which is immersed with a reference electrode in an aqueous conductive dispersion containing an enzyme and a non-ionic surfactant, the latter being present in an amount at least equal to the critical micelle concentration for the surfactant in the dispersion. A potential is then applied across the electrodes, causing the enzyme to deposit on the biosensor electrode. Next, the enzyme-deposited biosensor electrode is immersed in synthetic monomer and an electropolymerization procedure is carried out to create a polymer layer on the electrode which is intermingled with the initially deposited enzyme. In preferred forms, the electrode is next coated with a silane and then a polyurethane to complete the biosensor.
In more detail, the biosensor electrode (or more broadly an electrically conductive substrate) is typically formed of a metal alloy such as Ptxe2x80x94Ir or other noble metal; however, other conductive electrodes such as graphite electrodes can also be used. Normally, the electrode is in the form of a thin wire (having a diameter of from about 0.01 to 0.3 millimeters), and may have insulation over the majority of the length of the wire, leaving only a small stripped section to serve as an enzyme-receiving zone.
Virtually any enzyme can be deposited on the biosensor electrode, depending upon the nature of the analyte to be detected. The most useful enzymes are the oxidase enzymes, such as those selected from the group consisting of glucose, lactate, oxalate, D-aspartate, L-amino acid, D-amino acid, galactose, sarcosine, urate, ethanol, lysine, glutamate, cholesterol, glycerol, pyruvate, choline, ascorbate, and monoamine oxidases. Of these, the glucose, lactate, pyruvate, glutamate, cholesterol and choline oxidase enzymes are the most important.
During the initial deposition of the enzyme, it is preferred to employ an aqueous dispersion (normally a solution) which is buffered to a pH of about 7-8, more preferably about 7, containing the enzyme and a compatible surfactant (i.e., a surfactant which will facilitate enzyme deposition and not otherwise interfere with the desired electrodeposition). A variety of surfactants can be used in this context, although the nonionic surfactants are most preferred. One class of surfactants has been found to be particularly useful, namely the Triton surfactants, which are octylphenol ethoxylates, produced by the polymerization of octylphenol with ethylene oxide. A preferred surfactant is Triton X-100 which may be obtained from Sigma-Aldrich, Corp. (St. Louis, Mo.). The Triton X-100 product information sheet distributed by Sigma-Aldrich, Corp. is incorporated by reference herein. Triton X-100 comprises between about 9-10 moles of ethylene oxide per mole of octylphenol. Other surfactants of interest include nonylphenol ethoxylates, alkyl glucosides, alkyl maltosides, glucamides, alkyl polyoxyethylenes, alkyl glucopyranosides, alkyl thio-glucopyranosides, alkyl maltopyranosides, alkyl thio-maltoppyranosides, alkyl galacto-pyranosides, alkyl sucroses, glucamides, hyrdroxyethylglucamides, phenyl polyoxyethylenes, dimethylamine-N-oxides, cholate derivatives, n-octyl hydroxyalkylsulphoxides, sulphobetaines, and phosphocholine compounds.
An important feature of the enzyme deposition dispersion is that it contains the selected surfactant at a concentration of at least about the critical micelle concentration thereof, and more preferably a concentration within the range of from about the critical micelle concentration up to about 10 times the critical micelle concentration. Also, the relative amounts of surfactant and enzyme should be maintained. Generally, the molar ratio of the enzyme to the surfactant in the dispersion ranges from about 0.02 to 0.2, and more preferably from about 0.04 to 0.14.
The reference electrode used with the biosensor electrode during enzyme deposition is preferably AgCl/Ag. The potential applied across the electrodes during this process should be from about 1.1 to 1.4 volts versus the reference electrode. Such potential should be applied for a period of 40 to 80 minutes. The deposited enzyme on the electrode should have a thickness of from about 300 to 600 nm and more preferably from about 400 to 500 nm.
In the next step, a selected synthetic monomer is electropolymerized at the locale of the deposited enzyme so as to create a polymer layer which is intermingled with the enzyme. It is believed that the electropolymerization process causes the monomer to be oxidized at the surface of the electrode and encapsulates much of the enzyme. The thickness of the polymer layer formed is self-limiting, i.e., because the polymer is non-conducting, it cannot conduct electrons to become, in effect, an extension of the electrode. This means that the film will stop forming when communication with the electrode is interrupted.
The starting monomer for this step is selected with certain end properties in mind. Generally, the polymeric film should have a thickness and permeability consistent with the desired biosensor, but generally the polymeric layer should have a thickness of up to about 100 nm, and more preferably from about 10 to 100 nm. Second, the polymer film should have well-defined and reproducible permeability characteristics with optimal permeability for the enzyme substrate while excluding electroactive interferants such as ascorbate, urate and acetaminophen.
The single most preferred monomer for use in the electropolymerization step is phenol which should be present in an aqueous buffered phenol solution at pH 7. However, other candidate monomers include substituted and unsubstituted phenols. Preferred phenol monomers include 4-aminophenol, 1,4-dihydroxybenzene, 1,3-dihydroxybenzene, 1,2-dihydroxybenzene, 3-hydroxytyramine, 1,3,5-trihydroxybenzene, and 1,2,3-trihydroxybenzene.
In the electropolymerization procedure, the enzyme-deposited electrode is immersed in the monomer solution along with a reference electrode under an inert gas (e.g., argon) atmosphere. An appropriate potential is then applied to induce the electropolymerization reaction and create the desired polymer layer. In the next step of the preferred procedure, a silane film is applied over the polymer layer. This can be accomplished by dipping the electrode in a silane solution and applying a potential versus a reference electrode to enhance silane crosslinking. One suitable silane is (3-aminopropyl) trimethoxysilane, but other silanes such as 3-aminomethyl trimethoxysilane, 3-aminoethyl trimethoxysilane, 3-aminopropyl trimethoxysilane, 3-aminomethyl triethoxysilane, 3-aminoethyl triethoxysilane, and 3-aminopropyl triethoxysilane.
The next step in preparation of the sensor involves application of a polyurethane or other suitable coating over the silane film. This is done by dip-coating the electrodes with polyurethane solution in a suitable solvent and application of an appropriate potential versus the reference electrode. The final coating should have a thickness of from about 1 to 10 microns (1-10xc3x9710xe2x88x924 cm).
Once the sensor outer coating is applied, the sensor is conditioned. In the sensor conditioning process, the completed sensor is first dried at 4xc2x0 C. in a refrigerator for two or three days and is then transferred to a pH 7.4 phosphate buffered saline solution (0.05 M, containing 0.15 M NaCl). The buffer is changed every three days and the electrodes are conditioned continuously for two weeks with no potential applied to the electrodes. At the end of the two week period, the selectivity and sensitivity are measured and are hereafter referred to as the initial selectivity and initial sensitivity.
Sensors according to the invention should exhibit selectivity and sensitivity characteristics that are stable over a period of time. The selectivity characteristics of the polymer film should be stable for at least about 60 days, meaning that the selectivity should vary no more than xc2x110% relative to the initial selectivity over this time period. This feature of the polymer film is herein referred to as xe2x80x9cselectivity stability.xe2x80x9d The polymer film should also serve to stabilize the sensor response over time thus leading to stable glucose response or sensitivity. The sensitivity of the complete sensor should not be lower than 90% of the initial sensitivity and should not drift more than 0.5% per day for at least about 60 days. This stability of the sensor sensitivity is herein referred to as xe2x80x9csensitivity stability.xe2x80x9d
Finally, the linearity of the resulting sensor should be greater than 92% for a glucose sensor, calculated as follows:       %    ⁢          xe2x80x83        ⁢    linearity    =                    current        ⁢                  xe2x80x83                ⁢        response        ⁢                  xe2x80x83                ⁢        at        ⁢                  xe2x80x83                ⁢        25        ⁢                  xe2x80x83                ⁢        mM        ⁢                  xe2x80x83                ⁢        glucose        ⁢                  /                ⁢        25        ⁢                  xe2x80x83                ⁢        mM                    current        ⁢                  xe2x80x83                ⁢        response        ⁢                  xe2x80x83                ⁢        at        ⁢                  xe2x80x83                ⁢        5        ⁢                  xe2x80x83                ⁢        mM        ⁢                  xe2x80x83                ⁢        glucose        ⁢                  /                ⁢        5        ⁢                  xe2x80x83                ⁢        mM              xc3x97    100  